Method and apparatus for performing correction of transmission self-contamination in a medical imaging system

ABSTRACT

A method of acquiring transmission self-contamination data for correcting transmission scan data in a nuclear camera system including two transmission radiation sources and two scintillation detectors. A first source is for transmitting radiation to only a first detector and a second source is for transmitting radiation to only a second detector during a transmission scan of an object to be imaged. To acquire the calibration data, radiation is transmitted from the second source to the second detector while the first source is maintained in a non-transmitting state. While the second source is transmitting, the first detector is used to detect the radiation transmitted from the second source. A set of self-contamination data is then generated based on the radiation from the second source detected with the first detector. The calibration data is used to correct a subsequent transmission scan of the object by subtracting the self-contamination data from the transmission scan data.

FIELD OF THE INVENTION

The present invention pertains to the field of medical imaging. Moreparticularly, the present invention relates to attenuation correction innuclear medicine imaging systems.

BACKGROUND OF THE INVENTION

In the field of nuclear medicine, images of the internal structures orfunctions of a patient's body are formed by using gamma cameras todetect radiation emitted from within the body after the patient has beeninjected with a radiopharmaceutical substance. A computer systemgenerally controls the gamma cameras to acquire data and then processesthe acquired data to generate the images. Nuclear medicine imagingtechniques include Single-Photon Emission Computed Tomography (SPECT)and Positron Emission Tomography (PET). SPECT imaging is based on thedetection of individual gamma rays emitted from the body, while PETimaging is based on the detection of gamma ray pairs that are emitted incoincidence in opposite directions due to electron-positronannihilations. PET imaging is therefore often referred to as"coincidence" imaging.

One factor which has a significant impact on image quality in nuclearmedicine is non-uniform attenuation. Non-uniform attenuation refers tothe attenuation of radiation emitted from an organ of interest beforethe radiation can be detected. Such attenuation tends to degrade imagequality. A technique which has been used to correct for non-uniformattenuation is transmission scanning, in which gamma radiation from atransmission source is transmitted through the patient to acorresponding scintillation detector and used to form a transmissionimage. The transmission images provide an indication of the amountattenuation caused by various structures of the body and can thereforebe used to correct for attenuation in the emission images.

Many gamma camera systems include multiple detectors, multipletransmission sources, or both. In such systems, a particulartransmission source may be directed at a particular "target" detector atany given point in time. One problem which may be encountered with suchsystems is that of adequately shielding a transmission source from thedetectors which it is not supposed to be illuminating. For example, thecorrect interpretation of detected transmission radiation may dependupon knowing the location of the source of that radiation. Radiationdetected by a detector that is not the intended target is referred to astransmission self-contamination. Transmission self-contaminationintroduces inaccuracies into the transmission scan and thereforedegrades the quality of the emission images that are ultimatelycorrected by the transmission data. Even with proper shielding, it maybe impossible to completely prevent transmission self-contamination.Hence, it is desirable to have a technique for characterizing andcorrecting for transmission self-contamination in a nuclear medicineimaging system.

SUMMARY OF THE INVENTION

The present invention includes a method of performing a transmissionscan in a medical imaging system. The medical imaging system includes afirst source and a second source for transmitting radiation and a firstdetector and a second detector for detecting radiation. The first sourceis for transmitting radiation to the first detector but not to thesecond detector. The method includes the steps of using the first andsecond sources and the first and second detectors to perform atransmission scan of an object to acquire a set of transmission data,and using a second set of data to correct the set of transmission data.The second set of data is based on a calibration scan, which ischaracterized by a measurement of radiation transmitted by the firstsource which is detected by the second detector.

Another aspect of the present invention is a method of acquiringcalibration data for a nuclear camera system which includes a firstsource and a second source for transmitting radiation and a firstdetector and a second detector for detecting radiation. The methodincludes the steps of transmitting radiation from the second source tothe second detector, wherein the second source is for transmittingradiation to the second detector but not the first detector, and usingthe first detector to detect the radiation transmitted from the secondsource. A first set of data is then generated based on the radiationdetected with the first detector. The first set of data is forcorrecting a second set of data corresponding to an object to be imaged.

Other features of the present invention will be apparent from theaccompanying drawings and from the detailed description which follows.

BRIEF DESCRIPTION OF THE DRAWINGS

The present invention is illustrated by way of example and notlimitation in the figures of the accompanying drawings, in which likereferences indicate similar elements and in which:

FIG. 1 is a block diagram of a dual PET/SPECT gamma camera system.

FIG. 2 is a perspective view of the gantry and detectors of the gammacamera system of FIG. 1.

FIG. 3 is a block diagram of the computer system of the gamma camerasystem of FIG. 1.

FIG. 4 illustrates a coincidence timing circuit for switching betweenPET and SPECT modes.

FIG. 5A illustrates two radiation point sources illuminating twocorresponding detector imaging surfaces using fanbeam illuminationprofiles.

FIG. 5B illustrates two axially offset radiation point sourcesilluminating two corresponding detector imaging surfaces using fanbeamillumination profiles.

FIG. 6 illustrates an example of a sinogram which may be obtained usingthe configuration of FIGS. 5A with only one detector operating.

FIG. 7 illustrates an axial view of radiation point sources anddetectors in relation to the fields of view of the detectors.

FIG. 8A illustrates a radiation point source assembly.

FIG. 8B illustrates a radiation point source assembly according to anembodiment using a rotating aperture.

FIGS. 9A and 9B are flow diagrams illustrating overall routines forgenerating attenuation corrected PET images according to two differentembodiments.

FIG. 10 is a flow diagram illustrating a routine for performing atransmission scan of an object.

FIG. 11A illustrates countrate as function of position along thedetector imaging surface and the spatial windowing for transmission(T+E) and contamination (E) events.

FIG. 11B illustrates a relationship between emission and transmissionphotopeaks for one embodiment.

FIG. 12A is a block diagram illustrating a technique for correcting foremission contamination in a transmission image.

FIG. 12B is a flow diagram illustrating a routine for performingon-the-fly correction of emission contamination during a transmissionscan.

FIG. 13 is a graph illustrating effects of deadtime losses.

FIG. 14A is a block diagram illustrating a technique for performingon-the-fly correction of emission contamination and deadtime during atransmission scan.

FIG. 14B is a flow diagram illustrating a routine for performingon-the-fly deadtime correction and emission contamination and deadtimeduring a transmission scan.

FIG. 14C is a block diagram illustrating an alternative technique forperforming on-the-fly correction of emission contamination and deadtimeduring a transmission scan.

FIG. 15 illustrates a technique for rebinning singles events into arandoms sinogram.

FIG. 16 is a flow diagram illustrating a routine for correcting forrandoms during coincidence imaging.

DETAILED DESCRIPTION

A method and apparatus for measuring and correcting for transmissionself-contamination in a medical imaging system are described. In thefollowing description, for purposes of explanation, numerous specificdetails are set forth in order to provide a thorough understanding ofthe present invention. It will be evident, however, to one skilled inthe art that the present invention may be practiced without thesespecific details. In other instances, well-known structures and devicesare shown in block diagram form in order to facilitate description ofthe present invention.

Certain aspects of the present invention relate to attenuationcorrection of PET images in a dual-detector gamma camera system capableof both SPECT and PET imaging. An example of such a system isillustrated in FIG. 1 in block diagram form. A gamma camera systemhaving dual SPECT/PET capability is described in detail in U.S. Pat. No.5,585,637 and U.S. Pat. No. 5,608,221, both issued to Bertelsen et al.Certain aspects of such capability are described below as backgroundinformation. In addition, although the following description relates toan embodiment of the present invention having two gamma cameradetectors, the present invention is not limited to a two-detectorsystem.

The dual-mode gamma camera system 1 of FIG. 1 includes a processingsystem 20 coupled to a pair of scintillation detectors 10 and 11. Thedetectors 10 and 11 have imaging surfaces 12 and 13, respectively. Thedetectors 10 and 11 are mounted on a gantry 14, which can rotate thedetectors 10 and 11 either individually or in unison about an axis ofrotation, z (the "z axis"), which is perpendicular to the x-y plane. Apatient 15 to be imaged rests on a table 16 between the detectors 10 and11. The detectors are shown configured in a 180 degree orientation(i.e., offset 180 degrees relative to each other about the axis ofrotation) to facilitate coincidence (PET) imaging. Generally, theprocessing system 20 controls the gantry 14 to provide movement of thedetectors 10 and 11, controls the mode (PET vs. SPECT) of the detectors10 and 11, receives data acquired by the detectors 10 and 11, andgenerates images from that data. Each of the detectors 10 and 11includes a scintillation crystal, an array of photomultiplier tubes(PMTs) arranged in a conventional two dimensional matrix, and variousprocessing circuitry. Gamma camera detectors such as detectors 10 and 11are well-known in the art; accordingly a detailed description of theinternal components of detectors 10 and 11 and their operation is notnecessary to an understanding of the present invention and is notprovided herein. The scintillation crystals can be composed of sodiumiodine (NaI) and may be located between a collimator (not shown) and thePMT array.

The processing system 20 includes a programmable coincidence timingcircuit (CTC) 23 coupled to the detectors 10 and 11 and coupled to acomputer system 22. Note that in other embodiments, a CTC 23 may beincluded in one or both of the detectors 10 and 11. The computer system22 may include a conventional, general purpose workstation, asingle-board computer, or personal computer (PC). A signal 26 from thecomputer system 22 indicates to the CTC 23 the current mode of operation(i.e., SPECT or PET). Upon detection of a scintillation event in eitherdetector 10 or 11, lines 27 and 28, respectively, carry trigger pulsesto CTC 23. CTC unit 23 then generates valid event trigger signals overlines 17 and 18 for the detectors 10 and 11, respectively, according tothe selected mode of operation (SPECT or PET). The valid event triggersignals 17 and 18 are used by the detectors 10 and 11 to start (orreset) their accumulators (integrators), which accumulate (integrate)the energy of detected scintillation events and are therefore called"valid event" trigger signals. In the PET mode, integration is notstarted until a coincidence is detected between detector 10 and 11. InSPECT mode, an integration is started for each detector upon a triggerevent, regardless of coincidence. After integration and centroiding, thedetectors 10 and 11 output over lines 24 and 25, respectively, X and Yposition values and Z energy values.

FIG. 2 illustrates a perspective view of the gantry 14 and detectors 10and 11, according to one embodiment. FIG. 2 also shows two transmissionsource assemblies 30 and 31 for performing transmission scans. As willbe described below, each of the transmission source assemblies 30 and 31includes a single-photon ("singles") mode radiation source used toperform transmission scans for correcting PET images for the effects ofattenuation. In one embodiment, each of the source assemblies 30 and 31includes a Cs¹³⁷ point source. As will be described below, each of thesource assemblies 30 and 31 includes appropriate shielding as well ascollimation designed to provide a specific illumination profile. Sourceassembly 30 is mounted to a track assembly 32 adjacent to detector 10and outside the field of view (FOV) of detector 10. Source assembly 30has an aperture and is mounted so that the aperture faces detector 11 toallow radiation from source assembly 30 to illuminate detector 11.Similarly, source assembly 31 is mounted to a track assembly 33 adjacentto detector 11 and outside the FOV of detector 11. Source assembly 31further has an aperture and is mounted so that the aperture facesdetector 10 to allow radiation from source assembly 31 to illuminatedetector 10. Track assemblies 32 and 33 provide a mechanism fortranslating the source assemblies along the z axis in one embodiment ofthe present invention. Track assemblies 32 and 33 are rotatable aboutthe z axis in unison with detectors 10 and 11; consequently, the pointsources 30 and 31 at all times remain fixed relative to detectors 10 and11 in terms of their angular positions about the z axis.

FIG. 3 illustrates a block diagram of the computer system 22. Thecomputer system 22 comprises one or more central processing units (CPU)41, a read-only memory (ROM) 42, and a random access memory (RAM) 43,each coupled to a bus 52 for communicating information within the system22. Note that the bus 52 may comprise multiple physical buses coupledtogether by various bridges, controllers, and/or adapters. Also coupledto the bus 52 are a mass storage device 47, such as a magnetic oroptical disk and disk drive; a display device 45, such as a cathode raytube (CRT) or liquid crystal display (LCD); an alphanumeric keyboard 44;a pointing device 49, such as a mouse, trackball, or touchpad; and, acommunication device 46. The communication device 46 includes a highspeed communication port for communicating with the gantry 14 anddetectors 10 and 11 via signals 19, 24 and 25.

The computer system 22 executes software instructions to implementprocedures according to the present invention and various otherfunctions. Specifically, the CPU 41 may be configured to perform certainsteps in accordance with the present invention by software instructionsstored in RAM 43, ROM 42, mass storage device 47, or a combination ofthese devices.

As noted above, the CTC 23 (see FIG. 1) is used to control theoperational mode (SPECT or PET) of the gamma camera system 1. FIG. 4illustrates one embodiment of the CTC 23 in greater detail. FIG. 4 showsfour signals, 51, 52, 55, and 56. Signals 51 and 55 are trigger signalsgenerated by detector 10 in response to scintillation events and areprovided to CTC 23 over line 27. Signal 55 is generated by SPECTdetection electronics in detector 10, and signal 51 is generated by PETdetection electronics in detector 10. Signals 52 and 56 are triggersignals generated by detector 11 in response to scintillation events andare provided to CTC 23 over line 28. Signal 56 is generated by SPECTdetection electronics in detector 11 while signal 52 is generated by PETdetection electronics in detectors 11. Signals 51 and 52 from the PETdetection electronics of detectors 10 and 11, respectively, are providedas inputs to an AND gate 53. AND gate 53 outputs a signal 54, which isasserted only if signals 51 and 52 are in coincidence (i.e., bothasserted within a predetermined time window). The CTC 23 also includestwo double-input multiplexors 57 and 58. Multiplexor 57 receives asinput signal 55 from the SPECT detection electronics of detector 10 andsignal 54 from AND gate 53. Multiplexor 58 receives as input signal 56from the SPECT detection electronics of detector 11 and signal 54 fromAND gate 53.

A mode selection control signal 26 is coupled to the select inputs ofmultiplexors 57 and 58. The control signal 26 is used to switch betweenPET and SPECT modes of operation. The control signal 26 may result froma command entered by a user through a user interface provided by thecomputer system 22. When the control signal 26 has a value indicatingPET mode is desired, then an asserted signal over line 54 passes overboth line 17 to detector 10 and over line 18 to detector 11 as validevent trigger signals. When the control signal 26 has a value indicatingSPECT mode is desired, then the signal over line 55 is carried over line17 to detector 10, and the signal over line 56 is carried over line 18to detector 11 as valid event trigger signals. Signals over line 17 areused to trigger event integrators in the detection circuitry of detector10, and signals over line 18 are used to trigger event integrators inthe detection circuitry of detector 11.

Singles-Mode Fanbeam Transmission Scan

The present invention includes the use of two singles-mode radiationpoint sources to perform a transmission scan for purposes of performingattenuation correction on coincidence (PET) data. In one embodiment,each of the point sources is a Cs¹³⁷ source having an energy peak at 662keV. In the embodiment of FIG. 2, source assemblies 30 and 31 aremounted on the same side of the detectors 10 and 11 in the transaxial(x) direction. FIGS. 5A and 5B illustrate an alternative embodiment inwhich the source assemblies 30 and 31 are mounted on opposite sides ofdetectors 10 and 11 in the transaxial direction. Referring now to FIG.5A, source assemblies 30 and 31 (not shown) contain Cs¹³⁷ point sources30A and 31A, respectively. Point source 30A is mounted adjacent to theimaging surface 12 of detector 10, while point source 31A is mountedadjacent to the imaging surface 13 of detector 11. As indicated above,the point sources 30A and 31A remain fixed relative to detectors 10 and11 in terms of their angular positions about the z axis.

Transmission detection (spatial) windows 60 and 61 are defined on theimaging surfaces 12 and 13 of detectors 10 and 11, respectively, fordetecting transmission radiation transmitted from sources 31A and 30A,respectively. The transmission detection windows 60 and 61 are definedelectronically by the detectors and/or the computer system 22 in amanner that is well-known in the art. The transmission detection windows60 and 61 are defined for detection of photons within an energy rangecentered at 662 keV. Transmission detection window 60 is aligned withsource 31A relative to the z axis ("axially"), and transmissiondetection windows 61 is defined so that it is aligned axially withsource 30A. In the embodiment of FIG. 5A, the point sources 30A and 31Aare offset axially by an amount that is small relative to the axialfield of view (FOV) of detectors 10 and 51, such that transmissiondetection windows 60 and 61 are substantially adjacent in the zdirection. This offset reduces transmission self-contamination (i.e.,the undesirable detection of transmission radiation by the detectornearest to the transmitting source), which is discussed further below.

Also defined on the imaging surfaces 12 and 13 are emissioncontamination detection windows 62 and 63, respectively. The emissioncontamination detection windows 62 and 63 are used for purposes ofcorrecting for emission contamination of the transmission scan, as willbe described below.

During transmission scanning, the point sources 30A and 31A are scannedsynchronously across the FOVs of detectors 10 and 11 along the z axis.Further, transmission detection windows 60 and 61 are scannedsynchronously with their corresponding point sources, 31A and 30A. Onlyradiation detected within transmission detection windows 60 and 61 isrecognized for purposes of acquiring a transmission data set (i.e.,projection). In addition, each of the source assemblies 30 and 31 (seein FIG. 2) includes collimation designed to produce a fanbeamillumination profile, as shown in FIG. 5A, to substantially limittransmission radiation to transmission detection windows 60 and 61,respectively.

It may be desirable to have a greater axial offset between the pointsources 30A and 31A than that shown in FIG. 5A. Accordingly, FIG. 5Billustrates an alternative embodiment having a greater offset. A largeraxial offset between the point sources 30A and 31A may further reducetransmission self-contamination as well as cross-scatter of transmissionradiation into the wrong detection window.

During a PET imaging session, the detectors 10 and 11 are arranged in a180° orientation, and are used to detect emission radiation from anumber of angular positions about the z axis. Accordingly, at each ofthese angular positions about the z axis, a transmission scan isperformed by scanning the transmission radiation fanbeams and thecorresponding transmission detection windows 60 and 61 axially acrossthe FOVs of the detectors 10 and 11. Scanning of the fanbeams may beaccomplished by translating the source assemblies 30 and 31 axially.

In an alternative embodiment, scanning can be performed by maintainingthe point sources in a fixed position along the z axis and using arotating aperture to scan the fanbeam across the FOV of the opposingdetector. In such an embodiment, the count density at the correspondingdetector will vary depending upon the inclination angle of the fanbeam.For example, the count density will tend to decrease as the distancebetween the source and the illuminated portion of the detector increasesand, consequently, the illuminated area on the detector increases.However, assuming the camera system is calibrated using a blanktransmission scan (i.e., no object in the field of view), as isconventional practice in nuclear imaging, these effects will benormalized out when the actual images are generated.

In yet another embodiment, the transmission scan may be performed bymaintaining fixed axial (z) positions of the sources 30A and 31A whileeach of the sources 30A and 31A illuminates the entire imaging surfaceof the corresponding detector, rather than scanning a radiation beamacross the imaging surface. In this embodiment, a single-slice rebinningalgorithm may be adequate if the effective axial field of view issufficiently small, such that the incident angle at which thetransmission radiation impinges upon the detector surface is close to 90degrees. However, if the axial field of view to be scanned is relativelylarge, such that the incident angle becomes more acute, it may bedesirable to use a three-dimensional rebinning algorithm. An example ofa three-dimensional rebinning algorithm which may be used is the Fourierrebinning technique, which is described by M. Defrise et al., "Exact andApproximate Rebinning Algorithms for 3-D PET Data," IEEE Transactions onMedical Imaging, vol. 16, No. 2, April 1997.

The geometry illustrated in FIGS. 5A results in a base pattern sinogramsimilar to that illustrated in FIG. 6 (for one detector) for 32 angularstops over 360° of rotation of the detectors and point sources. In FIG.6, the diagonal lines regions represent regions of sinogram space inwhich data is acquired, while the gaps between the lines representregions in which there is no coverage. This effect occurs when thesinogram is generated by taking all of the possible 1024 differenttransverse positions that a detector can produce and then calculatingthe corresponding r and θ. Such gaps in the sinogram may produceartifacts in the reconstructed images. Therefore, it may be desirablefor the rebinning software to circumvent this effect. A solution is tostore raw detector coordinates and, during post-processing, fill in thesinogram by searching for and interpolating amongst the possible rawimage (projection) locations that could have contributed to a particularpoint in the sinogram.

FIG. 7 illustrates the geometry associated with the transmissionscanning of the present invention in terms of the FOVs of the detectors10 and 11 for one embodiment. Specifically, FIG. 7 illustrates a view ina transverse (x-y) plane according to an embodiment in which the sources30A and 31A are on the same side of detectors 10 and 11 in thetransaxial (x) direction. In FIG. 7, septa 74 and 75 are disposed alongthe imaging surfaces 12 and 13 of detectors 10 and 11, respectively.Point sources 30A and 31A are mounted outside the FOVs of detectors 10and 11. Such mounting avoids blocking the detectors and reducestransmission self-contamination. As noted above, a transmission scanacross the entire axial width of detectors 10 and 11 is performed ateach angular stop about the z axis. The aggregate effect of thesetransmission scans with the illustrated placement of point sources is atransmission FOV (in each transverse slice) represented by circle 70.The emission field of view (in each transverse slice) is represented bycircle 72.

In one embodiment, the point sources 30A and 31A are mounted outside theFOVs of detectors 10 and 11, such that the detectors themselves limitthe allowable beamwidth (in the transverse plane) of the fanbeamsgenerated by sources 30A and 31A. In such an embodiment, thetransmission FOV 70 is defined by two boundaries, an outside boundaryand an inside boundary. The outside boundary is defined by the outeredges of the transmission fanbeams 68 and 69 at each of the angularstops about the z axis, while the inside boundary is defined by thecircumference of circle 76. Thus, circle 76 represents a gap, or blindspot, in the transmission field of view 70. In order to prevent this gapfrom resulting in incomplete data acquisition, the computer system 22causes the table 16 (FIG. 1) to move vertically and horizontallyrelative to the z axis in dependence on the angular positions of thedetectors 10 and 11 about the z axis in order to provide full coverageof the object of interest. Such table motion effectively increases thetransmission FOV 70. In one embodiment, table motion is controlled bythe gantry 14, which includes a dedicated microprocessor (not shown). Atechnique for providing table motion in a medical imaging system isdescribed in U.S. Pat. No. 5,444,252 of Hug et al., which is assigned tothe assignee of the present invention.

FIG. 8A illustrates the point source assembly 30 according to oneembodiment. It should be noted that point source assembly 31 issubstantially identical to point source assembly 30 unless specificallystated otherwise. The assembly 30 includes a number of lead shieldingstructures 81, which partially enclose the point source 30A. The leadstructures 81 are encased by aluminum or steel brackets 82 and 83.Bracket 83 forms a means for mounting source assembly 30 to trackassembly 32 (see FIG. 2) to allow axial translation of source assembly30.

FIG. 8B shows an alternative embodiment of a point source assembly oruse in an embodiment in which the point sources are not translated alongthe z axis in order to provide the scanning of the fanbeam. The sourceassembly 30A1 is fixed axially but includes an aperture 82 that isrotated about an axis 84 to provide the scanning of the fanbeam. Thesource assembly 30A1 is mounted appropriately to the gantry 14 with axis84 is parallel to the x axis. The point source in this embodiment (notshown) is encased with appropriate shielding 81 similar to that ofassembly 30A in FIG. 8A. The casing 81 and aperture 82 are rotated abouta shaft 83 (i.e., about axis 84).

FIG. 9A illustrates an overall routine according to one embodiment foracquiring both coincidence (emission) data for a PET study, as well assingles transmission data for attenuation correction of the coincidencedata. In step 901, the initial study parameters are set up in thecomputer system 22. These parameters include, for example, the totalnumber of angular stops about the z axis and the total acquisition timeat each stop for both the emission scan and transmission scan. Next, instep 902, the detectors 10 and 11 are configured in a 180° orientationabout the z axis to enable coincidence detection. In step 903 thedetectors 10 and 11 are rotated (together with source assemblies 30 and31) to the first (or next) angular stop about the z axis. In step 904,the detectors are configured for detection of emission data incoincidence mode. After acquiring coincidence data for the prescribedtime period in step 905, the detectors are then configured for detectionin the singles mode in step 906. Next, in step 907, a transmission scanis performed in the manner described above (i.e., by scanning thetransmission radiation fanbeams across the detector imaging surfaces),and the transmission data is acquired as singles data. If there areadditional angular stops at which data is to be acquired (step 908),then the routine repeats from step 903. Otherwise, in step 909 theemission image is reconstructed and corrected for attenuation using thetransmission image data.

FIG. 9B illustrates another overall routine for acquiring bothcoincidence (emission) data for a PET study and transmission data forattenuation correction, according to a second embodiment. In theembodiment of FIG. 9B, the emission data is acquired first followed byacquisition of transmission data. More specifically, the studyparameters are set up in step 921, and the detectors are configured in a180 degree orientation in step 922. Next, after configuring thedetectors for coincidence mode acquisition in step 923, emission data isacquired for the complete range of projection angles, rotating thedetectors between angular stops about the z axis as required (steps 924,925 and 926). After the emission data is acquired, the detectors arereconfigured for singles-mode acquisition in step 927, and transmissiondata is acquired for the complete range of projection angles, rotatingthe detectors between angular stops about the z axis as required (steps928, 929 and 930). An emission image is then reconstructed and correctedfor attenuation using the transmission image data in step 931.

Correction of Emission Contamination in the Transmission Scan

If the transmission scan is performed after injection of theradionuclide into the patient, emission activity will be present duringthe transmission scan. Accordingly, some of the emission activity may bedetected (undesirably) within transmission detection windows 60 and 61(see FIGS. 5A and 5B). The effect of this emission contamination in thetransmission detection scan is illustrated in FIG. 11A, which plotscountrate as a function of axial (z) position along the imaging surface12 of a detector. A baseline countrate 88 of emission activity existsacross the entire imaging surface 12 of the detector. In addition,within the transmission detection window 60, there is additionalcountrate attributable to transmission radiation from the correspondingtransmission source 31A (not shown in FIG. 11A). Energy discriminationas a means for distinguishing between emission activity and transmissionactivity becomes relatively ineffective if the emission source andtransmission source have photopeaks that are close together (e.g., aCs¹³⁷ singles transmission source with a photopeak at 662 keV and aFlouro Deoxy Glucose, or FDG, coincidence emission source with aphotopeak at 511 keV). The reason energy discrimination becomesrelatively ineffective is that, as illustrated in FIG. 11B, a portion ofthe emission energy distribution 87 will fall into the transmissionenergy acceptance range 89. Note that the energy acceptance range is notto be confused with transmission detection windows 60 and 61, which arespatial windows. As a result, some of the emission activity impinging onthe transmission detection windows 60 and 61 will be incorrectlydetected as transmission activity, thereby introducing inaccuracy intothe transmission image.

Consequently, the present invention includes a technique for reducingemission contamination in the transmission scan. Specifically, emissioncontamination detection windows 62 and 63 are defined on the imagingsurfaces 12 and 13 of detectors 10 and 11, respectively, for detectionof photons in an energy acceptance range centered at the photopeak ofthe transmission source (i.e., 662 keV if a Cs¹³⁷ transmission source isused). That is, both the transmission detection windows 60 and 61 andthe emission contamination detection windows 62 and 63 have energyacceptance ranges centered at the transmission photopeak, as shown inFIG. 11B.

Because the transmission detection windows 60 and 61 will receive someemission activity, those windows will henceforth be referred to in thisdescription as the "T+E" (Transmission+Emission) windows 60 and 61 tofacilitate description. Similarly, the emission contamination detectionwindows 62 and 63 (see FIGS. 5A, 5B, and 11) will henceforth be referredto as the "E" windows. Note that the emission singles rate fallingwithin the transmission energy acceptance window does not changesubstantially over the short acquisition time period, regardless ofwhether the transmission beam is over the T+E window 60. Therefore, theemission count detected within the E window 62 can be used to subtractout the emission contamination within the T+E window 60. FIG. 10illustrates in greater detail the step 907 (FIG. 9A) or 929 (FIG. 9B) ofacquiring transmission data according to the above-mentioned technique.

At each angular stop, the transmission scan is characterized by threeconcurrent steps, 1001A, 1001B, and 1001C. In step 1001A, thetransmission sources and the T+E detection windows 60 and 61 are scannedaxially across the FOVs of the detectors 10 and 11, respectively. As theT+E windows 60 and 61 are scanned, transmission data is acquired in step1001B based on radiation detected within the T+E windows 60 and 61 andrebinned into sinograms using an axial rebinning algorithm in step1001C. This process repeats until the entire axial FOVs of detectors 10and 11 have been scanned (step 1002), at which time the sourceassemblies 30 and 31 are returned to their initial axial positions instep 1003. Alternatively, after step 1002 the source assemblies 30 and31 can be left at their final axial position and then scanned in theopposite direction axially for the next angular position of thedetectors 10 and 11.

The present invention provides for a correction of emissioncontamination in the transmission scan by measuring the emissioncountrate at various positions and points in time and correcting foremission contamination on-the-fly (on an event-by-event basis) as thetransmission scan is performed, as will be described in greater detailbelow. This technique includes defining the E windows 62 and 63 on theimaging surfaces 12 and 13 of detectors 10 and 11, respectively. The Ewindows 62 and 63 are scanned in synchronization with, but are offsetaxially from; the transmission T+E windows 60 and 61. For a given axialposition of the E windows 62 and 63, the number of counts detected inthe E windows 62 and 63 provides a good approximation of the emissionactivity impinging on the T+E windows 60 and 61 when the T+E windows 60and 61 are located at the same position. Consequently, in accordancewith one embodiment of the present invention, each time a count isdetected in an E window of a detector, a count is removed from thecorresponding location in a transmission projection buffer representingthe data acquired in the T+E windows 60 and 61. The result is toeffectively remove virtually all of the emission contamination from thetransmission image.

Note that the fanbeam collimation of the point sources 30A and 31Aenables this technique to be performed in conjunction with a singlestransmission source. More specifically, the fanbeam collimation enablesthe simultaneous acquisition of both emission and singles transmissiondata with the same detector and on-the-fly correction. This technique isin contrast with prior techniques which make use of uncollimatedcoincidence sources and/or do not perform on-the-fly correction ofemission contamination. See, e.g., R. J. Smith et al., "SimultaneousPost Injection Transmission and Emission Contamination Scans in a VolumeImaging PET scanner, " 1995 IEEE Nuclear Science Symposium and MedicalImaging Conference Record, Vol. 3, pages 1781-85, 1995, and R. J. Smithet al.,"Post Injection Transmission Scanning in a Volume Imaging PETCamera," IEEE Transactions on Nuclear Science, vol. 41, No. 4, August1994.

FIG. 12A further illustrates the above-described technique forcorrecting for emission contamination in the transmission scan.Transmission and emission radiation is detected by the detectors 10 and11 using the scanning T+E and E detection windows, as described above.Information on the detected events (i.e., X and Y position and energylevel Z) is provided to an axial rebinning algorithm 98, which receivesas input the axial (z) positions of the sources 30A and 31A and theangular positions of the detectors and sources about the z axis. Foreach count detected in either the T+E window or the E window of one ofthe detectors 10 and 11, if the count was detected in an T+E window(block 99), then one count is added to the corresponding location in thetransmission projection buffer 100, and if the count was detected withinan E window, then one count is subtracted from the correspondinglocation in the transmission projection buffer 100. The countssubtracted from the transmission projection buffer 100 willsubstantially equate to the emission contamination counts undesirablyadded to the corresponding locations in the projection buffer 100. Notethat the inputting of the source positions to axial rebinning algorithm98 allows counts to be added or subtracted from the proper locations ofthe transmission projection buffer 100, such that correction of emissioncontamination can be performed on an event-by-event basis and in aspatially dependent manner.

FIG. 12B illustrates the step 907 (or 929) of performing thetransmission scan according to an embodiment which uses emissioncontamination correction in accordance with FIG. 12A. The routine ischaracterized by three concurrent processing paths. In the first path,it is determined in step 1201A whether an event has been detected in aT+E window. If so, the corresponding location in the transmissionprojection buffer 100 is incremented by one count in step 1202A; if not,the processing path repeats from the beginning. In the second processingpath, the transmission sources, T+E windows, and E windows are scannedaxially across the FOVs of their respective detectors to acquire data instep 1201B. In the third processing path, it is determined in step 1201Cwhether an event has been detected in an E window. If so, thecorresponding location in the transmission projection buffer 100 isdecremented by one count in step 1202B; if not, the processing pathrepeats from the beginning. After performing either step 1202A, 1201B,or 1202B, then if the entire axial FOV has not been scanned, then theroutine repeats from the beginning; otherwise, the routine ends.

Note that the width of the T+E and E windows can be varied based onaxial position to achieve the best windowing effect, such as when awindow reaches the edge of the imaging surface. However, for any givenaxial position, the width of the T+E and E windows remains constant toensure accurate correction.

Deadtime Correction

One problem associated with conventional gamma cameras is deadtime loss.Deadtime refers to the inability of a scintillation detector todistinguish two distinct scintillation events which occur very closetogether in time. Deadtime loss can be defined as the difference betweenthe true countrate and the observed countrate which results fromdetector deadtime. FIG. 13 illustrates the effect of deadtime losses inthe form of a plot of observed countrate against true countrate. Line 91represents the ideal (but unrealistic) case in which there is nodeadtime loss; in that case, the observed countrate OC equals the truecountrate C1. In contrast, line 90 represents the response of a gammacamera system that is subject to deadtime loss; in that case, theobserved-countrate OC is lower than the true countrate C2. Note that thedeadtime loss is dependent upon the singles rate; that is, as thesingles rate (true countrate) increases the deadtime loss (differencebetween the true countrate and the observed countrate) also increases.

One known technique for correcting for deadtime loss is to apply asingle, global correction factor, which is not applied until after thedata has been acquired. See, e.g., R. J. Smith et al., "SimultaneousPost Injection Transmission and Emission Contamination Scans in a VolumeImaging PET scanner," 1995 IEEE Nuclear Science Symposium and MedicalImaging Conference Record, vol. 3, pages 1781-85, 1995. However, the useof a global correction factor does not account for the spatialdependency of deadtime losses. In particular, deadtime loss is dependentupon the singles rate, which is dependent upon both axial position andprojection angle. Therefore, the use of a global deadtime correctionfactor may introduce inaccuracies into the transmission image.

Hence, the present invention includes a technique for correcting fordeadtime losses which takes into account the spatial dependency ofdeadtime losses. More specifically, the present invention includes atechnique for correcting for deadtime losses on an event-by-event basis(on-the-fly). Because deadtime is dependent upon the singles rate, asingles rate vs. deadtime calibration curve can be derived for a givengamma camera system. Hence, in accordance with the present invention, a,deadtime versus singles rate calibration curve is empirically derivedfor the gamma camera system 1 and then used to create a look-up table ofdeadtime correction factors for various different singles rates. In oneembodiment, each correction factor in the look-up table is a factor bywhich an observed count is multiplied during an imaging session beforebeing added to the transmission projection, in order to compensate fordeadtime losses. Higher singles rates will correspond to higher deadtimelosses and, therefore, higher correction factors from the look-up table.The look-up table can be an integer map, which may be created to have anumber of entries chosen so as not to compromise the speed of rebinning.

As an example of this technique, if the current singles rate correspondsto no significant deadtime loss, then the corresponding location in thetransmission projection buffer can be increased by 50 counts, ratherthan one count, for each detected count. On the other hand, if there isa singles rate corresponding to a 2% deadtime loss, then thecorresponding location in the transmission projection buffer can beincreased by 51 counts rather than one count. Note that increasing thenumber of counts in this way introduces an artificial magnificationfactor in the transmission projection. The above example would introducea magnification factor of 50 into the transmission projection. Thismagnification can be removed during reconstruction, however, by scalingdown the image accordingly at that time. It should be noted that theaccuracy of the deadtime correction depends upon the magnificationfactor selected, the accuracy of the initial calibration, and theassumption that there is little or no variance of deadtime within thetransmission beam.

FIG. 14A illustrates a technique for implementing deadtime correctionon-the-fly, as described above. In one embodiment, this technique isimplemented in conjunction with the emission contamination correctiontechnique described above. Specifically, the process flow of FIG. 14A isessentially the same as that of FIG. 12A with the exception of theaddition of the singles rate deadtime look-up tables (LUTs) 101. If acount is detected in a T+E window (block 99), then rather than addingone count to the appropriate location in the transmission projectionbuffer 100, as in the case of FIG. 12A, X counts are added to thatlocation in the buffer, where X is determined from the singles ratedeadtime look-up table 101 based on the current singles rate. Similarly,if an event is detected in the E window, then rather than subtracting asingle count from the appropriate location in the transmissionprojection buffer, X counts are subtracted from that location, where Xis determined from the singles rate deadtime look-up table 101 based onthe current singles rate. Note that, in one embodiment, the singles rateused for this purpose is the global singles rate (i.e., the singles rateobserved across an entire detector) rather than the singles rate withineither a T+E window or an E window.

As with correction of emission contamination, the correction fordeadtime loss is performed on an event by event basis. Accordingly, thesource position (angular position about the z axis and axial position)is input to the axial rebinning algorithm 98 to enable each event to beassociated with the appropriate location in the transmission projectionbuffer 100. This technique is in contrast with the above-mentionedtechnique in which a single, global correction factor is applied, whichdoes not take into account the spatial dependency of deadtime losses.

FIG. 14B illustrates the step 907 (or 929) of performing thetransmission scan according to an embodiment which uses on-the-flycorrection of both deadtime loss and emission contamination, asdescribed above. The routine is characterized by four concurrentprocessing paths. In the first processing path, it is determined in step1401A whether an event has been detected in a T+E window; if not, theprocessing path repeats from the beginning. If an event has beendetected in the T+E window, then in step 1402A, the correspondinglocation in the transmission projection buffer 100 is incremented by Xcounts, where X is determined in another concurrent processing path, aswill be described below. A second processing path consists of step1401B, in which the transmission sources 30A and 31A, the T+E windows 60and 61, and the E windows 62 and 63 are scanned axially to acquire data.A third processing path begins with step 1401D, in which it isdetermined whether an event has been detected in an E window; if not,the processing path repeats from the beginning. If an event has beendetected in the E window, then in step 1402D the corresponding locationin the transmission projection buffer 100 is decremented by X counts,where X is determined in the fourth concurrent processing path asfollows. In the fourth concurrent processing path, the current singlesrate is computed in step 1401C. As noted above, the singles rate in thisembodiment is the global singles rate for the detectors 10 and 11 forthis particular detector angle. Based on the current singles rate, thecurrent count factor X is determined from the look-up table in step1402C. After completion of any of the four concurrent processing paths,then the routine repeats from the beginning if the entire axial field ofview has not yet been scanned (step 1403). Otherwise, the routine ends.

FIG. 14C shows an alternative embodiment for implementing on-the-flydeadtime correction and emission contamination correction. Note that theembodiment of FIG. 14C, as with other aspects of the present invention,can be implemented in software, hardware, or a combination thereof. Asin the embodiments described above, both emission contaminationcorrection and deadtime correction are performed on an event by eventbasis. In contrast with the embodiment of FIGS. 14A and 14B, however,each observed count is not multiplied by a factor in the embodiment ofFIG. 14C. Rather, each count generally is represented as one count inthe transmission projection buffer 100. However, additional counts maybe added or subtracted to appropriate locations of the transmissionprojection buffer 100 depending upon the observed singles rate at agiven point in time and the spatial window (T+E or E) in which a givencount is detected.

The technique of FIG. 14C includes a look-up table 105, which outputs avalue A that is based on the current singles rate. The technique alsoincludes counters 106 and 107, each of which counts up in response toits clock input to a maximum value MAXCOUNT and then automaticallyresets to zero. Counter 106 is clocked by signal E_(x),y, which isasserted each time an event is detected in the E window. Counter 106outputs a signal B1. Similarly, counter 107 receives a signal T+E_(x),yat its clock input, which is asserted each time an event is detected inthe T+E window, and outputs a signal B2.

The technique of FIG. 14C also includes two comparators 109 and 110.Comparator 109 receives as input signals A and B1 and asserts its outputsignal when B1 is greater than A (i.e., when the output of counter 106exceeds the output of look-up table 105). Comparator 110 receives asinput signals A and B2, and asserts its output when B2 is greater than A(i.e., when the output of counter 107 exceeds the output of look-uptable 105).

Look-up table 105 is generated in a manner similar to that describedabove in connection with FIGS. 14A and 14B. That is, look-up table 105is created based on empirical data and provides output signal A having avalue based on the global input singles countrate. Specifically, thevalue A output by look-up table 105 is based on the following equation:

    A=MAXCOUNT*(COUNTRATE.sub.observed +(COUNTRATE.sub.true -COUNTRATE.sub.observed)/2)

where:

COUNTRATE_(true) is the true countrate if the deadtime were zero,

COUNTRATE_(observed) is the intrinsic observed countrate, and

MAXCOUNT is the maximum value of the free running counters 106 and 107.

Each time an event is detected in the T+E window (i.e., each timeT+E_(x),y is asserted), block 112A causes one event to be added to theappropriate x,y location of transmission projection buffer 100. Inaddition, if B2 is greater than A, block 112B causes an additional eventto be added to the appropriate x,y location of transmission projectionbuffer 100 in response to assertion of signal T+E_(x),y. Adding suchadditional events compensates for deadtime losses. However, events mustalso be removed from transmission projection buffer 100 to correct foremission events detected in the T+E window. Accordingly, each-time anevent is detected in the E window (i.e., each time E_(x),y is asserted),block 111B causes one event to be subtracted from the correspondinglocation in the transmission projection buffer 100. In addition, if B1is greater than A, block 111A causes an additional event to besubtracted from the corresponding x,y location in the transmissionprojection buffer 100 in response to assertion of signal E_(x),y. Hence,the technique illustrated in FIG. 14C provides correction on anevent-by-event basis of both deadtime losses and emission contaminationin the transmission scan.

Note that the predetermined contents of the look-up table 105 as well asthe value of MAXCOUNT will determine the actual response of theillustrated embodiment (i.e., how frequently additional events will beadded or subtracted to/from the transmission projection for a givensingles rate). As in the embodiments discussed above, this technique isadvantageous in comparison to previous techniques which do correct forthe spatial variances in deadtime or emission contamination.

Transmission Self-Contamination

Another factor which can cause degradation of image quality isself-contamination of the transmission scan. Transmissionself-contamination refers to the detection of transmission radiation bya detector which is not the intended target of the source from which theradiation originates. More specifically, referring to FIG. 7,transmission self-contamination includes radiation transmitted by source30A that is unintentionally detected by detector 10 and radiationtransmitted by source 31A that is unintentionally detected by detector11. The present invention provides certain features to reduce suchcontamination.

Mounting the sources 30A and 31A outside the FOVs of the detectors 10and 11 serves to reduce the likelihood of transmissionself-contamination, as does proper shielding within source assemblies 30and 31, in which sources 30A and 31A are enclosed. A third technique forreducing transmission self-contamination is to quantify thatself-contamination and subtract it out from the transmission image. Thisform of calibration can be performed by performing a transmission scanin which only the source that is closest to the detector to becalibrated is transmitting. That source as well as the T+E and E windowson the detector to be calibrated are then scanned across the FOVs of thedetectors. Two images are generated for each detector based on thiscalibration scan, one image for the transmission self-contamination inthe T+E window, and one for the transmission self-contamination in the Ewindow. Note that if the translation of the transmission source isnon-uniform inside the FOV, it may also be necessary to correct for thedifferences in count densities that result from this effect. After thetwo images are acquired, they are low-pass filtered and stored in abuffer, such that they are accessible to the rebinning algorithm.

Correction of the transmission self-contamination can be performed bysubtracting the difference between the self-contamination calibrationprojections (i.e., T+E projection minus E projection) from thetransmission projections and then reconstructing the transmissionprojections.

Randoms Correction

Another problem encountered in PET imaging system is that of randomcoincidences ("randoms"). PET imaging systems produce images based uponthe detection of coincidence events. Coincidence events are normallyidentified as two events observed by opposing detectors which occurwithin a relatively narrow time window. However, some pairs of eventsdetected within this time window are not the result of true coincidences(i.e., based on a single positron-electron annihilation) but are basedon separate annihilation or Compton events. The misinterpretation ofsuch randoms as true coincidences contributes to inaccuracies in theimaging process and, therefore, degrades image quality.

One known technique for providing randoms correction is a directmeasurement technique in which a parallel time circuit for thecoincidence detection circuit is provided with a time delay on thesignal from one detector. The delay is made large enough so that no truecoincidence events are registered in the parallel time circuit. As aresult, events which are detected in the parallel time circuit canresult only from random coincidences, which can then be subtracted fromthose counted as true coincidences. One problem with this technique,however, is that it tends to be a relatively complex hardware solutionthat adds to the cost and size of the gamma camera system. In addition,since this solution is based on direct measurement of (random)coincidences, and coincidences represent only a small fraction of theoverall countrate, the resulting randoms data that is acquired tends tobe noisy. Another approach is to form an estimate of the randoms basedupon a mathematical model; however, because that approach is based uponan estimate, it is inherently subject to inaccuracies. For example, suchan estimate might not take into consideration the spatial variations ofrandoms. Hence, the present invention provides a technique forcorrecting for randoms which overcomes these disadvantages.

In general, randoms R can be calculated according to the equationR=S1*S2*W, where S1 represents the singles rate from one detector, S2represents the singles rate from the other detector, and W (also knownas 2τ) represents the duration of the coincidence time window. Inaccordance with the present invention, each of the detectors 10 and 11is treated as a collection of multiple subdetectors, each having asingles rate that is a subset of the overall singles rate (S1 or S2).For example, referring to FIG. 15, the total singles rate S1 fordetector 10 can be expressed as:

S1=S11+S12+S13+. . . +S1n.

Similarly, the total singles rate S2 for detector 11 can be expressedas:

S2=S21+S22+S23+. . . +S2n.

Therefore, the total randoms R can be expressed as: ##EQU1## where Ri(i=1, 2, . . . , (n*n)) is the randoms caused by any subcombination of asubdetector from detector 10 and a subdetector from detector 11. Thus,the above equation represents the concept that the global randoms aresimply the sum of all the randoms contributed by each subdetectorelement pair.

In accordance with the present invention, by applying the concept of theabove equation, a randoms sinogram can be obtained, which can be used tosubtract out the randoms from the emission data. FIG. 16 illustrates aroutine for performing randoms correction using this technique. In step1601, the study parameters (i.e., number of angular stops andacquisition time at each stop) are set up as described above. In step1602, detectors 10 and 11 are configured in a 180° degree orientationabout the z axis, and in step 1603 the detectors are configured foracquisition in singles-mode. Next, the detectors are rotated to thefirst (or next) angular stop in step 1604A. Concurrently with rotatingthe detectors to the next angular stop, emission data is acquired insingles-mode and rebinned into randoms sinograms in step 1604B. In oneembodiment, the data is acquired in step 1604B in the form of a numberof "snapshots" taken at various angular positions between two angularstops. After the detectors have been rotated to the next angular stopand the singles data have been acquired, the detectors are configuredfor acquisition in coincidence mode in step 1605. Next, coincidence datais acquired and rebinned into coincidence sinograms in step 1606 for theprescribed time interval. If emission data have been acquired at all ofthe angular stops (step 1607), then the routine proceeds to step 1608,in which the randoms sinograms are subtracted from the correspondingcoincidence sinograms. If there are additional angular stops, then theroutine repeats from step 1603, in which the detectors are temporarilyreconfigured for singles mode. After subtracting out the randoms, theemission image is reconstructed in step 1609.

FIG. 15 illustrates how the singles rate data is rebinned into a randomssinogram. Detector 10 is divided into n virtual subdetectors S11, S12, .. . S1n and detector 11 is divided into n virtual subdetectors S21, S22,. . . S2n. Counts detected by detectors 10 and 11 are rebinned intorandoms sinogram 92. For each detector angle about the z axis, therebinned data appears as a diamond in the randoms sinogram. The finalrandom sinogram is a summation of all the individual sinograms for eachaxial position. In one embodiment, the single-slice rebinning techniqueis used in the axial direction.

Note that the singles rate is much higher than the coincidence rate.Accordingly, a comparatively small period of time is required to acquirethe singles data for purposes of randoms correction than is required toacquire the total emissions data. For example, for a coincidencefraction equal to 5% of the singles rate, the acquisition of singleswould require only 1/20 of the total emission acquisition time. Thus,for a 20 minute total emission scan time, for example, only one minuteof total detector rotation time would be required to acquire singlesdata for randoms correction.

It will be recognized that the randoms correction technique of thepresent invention takes into consideration and corrects for spatialvariations in randoms by detecting the spatial variations in thesingles. However, it should also be noted that, because randomsgenerally do not yield high frequency data in comparison to truecoincidences, the randoms data can be stored in a matrix that is smallerthan the normal emission projection matrix. For example, a 32×32×48matrix may be used for randoms data as compared to a 128×128×96 matrixfor the normal emission projection. The randoms matrix can then beexpanded during correction.

It should also be noted that, in practice, calculating randoms R asS1*S2*W may not hold true. Therefore, it may be desirable to modify thisequation using a calibration factor, i.e., R=Cal(S1, S2)*S1*S2*W, whereCal(S1,S2) is a calibration factor that may or may not depend upon S1and S2. This calibration factor can be modeled as an empirical functionthat is derived using measurements with simple phantoms.

Thus, a method and apparatus for measuring and correcting fortransmission self-contamination in a medical imaging system have beendescribed. Although the present invention has been described withreference to specific exemplary embodiments, it will be evident thatvarious modifications and changes may be made to these embodimentswithout departing from the broader spirit and scope of the invention asset forth in the claims. Accordingly, the specification and drawings areto be regarded in an illustrative rather than a restrictive sense.

What is claimed is:
 1. A method of performing a transmission scan in amedical imaging system, the medical imaging system including a firstsource and a second source for transmitting radiation and a firstdetector and a second detector for detecting radiation, the methodcomprising the steps of:using the first and second sources and the firstand second detectors to perform a transmission scan of an object toacquire a set of transmission data, wherein the first source is fortransmitting radiation to the first detector and not to the seconddetector; and using a second set of data to correct the set oftransmission data, wherein the second set of data is based on acalibration scan, the calibration scan characterized by a measurement ofradiation transmitted by the first source which is detected by thesecond detector.
 2. A method according to claim 1, wherein the set oftransmission data comprises a plurality of counts, and wherein the stepof using the second set of data to correct the set of transmission datacomprises the step of subtracting counts from the set of transmissiondata based on the second set of data.
 3. A method of acquiringcalibration data for a nuclear camera system, the nuclear camera systemincluding a first source and a second source for transmitting radiationand a first detector and a second detector for detecting radiation, themethod comprising the steps of:transmitting radiation from the secondsource to the second detector, the second source for transmittingradiation to the second detector and not to the first detector; usingthe first detector to detect the radiation transmitted from the secondsource; and generating a first set of data based on the radiationdetected with the first detector, the first set of data for correcting asecond set of data corresponding to an object to be imaged.
 4. A methodaccording to claim 3, further comprising the step of maintaining thefirst radiation source in a non-transmitting state during the step oftransmitting radiation from the second source to the second detector. 5.A method according to claim 3, wherein the second set of data comprisestransmission data acquired from a transmission scan of the object usingthe first and second transmission sources and the first and the firstand second detectors.
 6. A method according to claim 3, furthercomprising the steps of:using the first and second sources and the firstand second detectors to perform a nuclear medicine transmission scan ofthe object to acquire the second set of data; and using the first set ofdata to correct the second set of data for transmissionself-contamination.
 7. A method according to claim 6, wherein the secondset of data comprises a plurality of counts representing detectedradiation, and wherein the step of using the first set of data tocorrect the second set of data for transmission self-contaminationcomprises the step of subtracting counts from the second set of databased on the first set of data.
 8. A method according to claim 3,wherein the step of transmitting radiation from the second source to thesecond detector comprises the step of transmitting radiation from thesecond source to,the second detector while maintaining the firstradiation source in a non-transmitting state.
 9. In a nuclear camerasystem, a method of performing a transmission scan of an object to beimaged, the nuclear camera system including a first radiation source anda second radiation source for transmitting radiation and a firstdetector and a second detector for detecting radiation, the methodcomprising the steps of:scanning radiation from the second radiationsource across an imaging surface of the second detector whilemaintaining the first radiation source in a non-transmitting state, thefirst radiation source for transmitting radiation to only the firstdetector, the second radiation source for transmitting radiation to onlythe second detector; detecting radiation from the second radiationsource with the first detector during the scanning step; generating afirst set of data based on the radiation detected within the firstdetector; using the first and second radiation sources and the first andsecond detectors to perform a transmission scan of the object to acquirea second set of data; and using the first set of data to correct thesecond set of data.
 10. A method of correcting for transmissionself-contamination of a transmission scan in a nuclear camera system,the nuclear camera system including a first radiation source fortransmitting radiation to a first scintillation detector and a secondradiation source for transmitting radiation to a second scintillationdetector, wherein the first and second scintillation detectors each havean imaging surface, the method comprising the steps of:defining atransmission detection window within a first portion of the imagingsurface of the first scintillation detector, the transmission window forreceiving transmission radiation from the first transmission source;defining an emission detection window within a second portion of theimaging surface of the first scintillation detector different from thefirst portion, the emission detection window for receiving radiationemitted from within an object to be imaged; scanning radiation from onlythe second radiation source across the imaging surface of the secondscintillation detector; scanning the transmission detection windowacross the imaging surface of the first scintillation detector duringthe step of scanning radiation from the second radiation source;scanning the emission detection window across the imaging surface of thefirst scintillation detector during the step of scanning radiation fromthe second radiation source; detecting radiation from the secondradiation source within the transmission detection window and theemission detection window during the step of scanning radiation from thesecond radiation source; and generating a set of transmissionself-contamination data based on the radiation detected within thetransmission detection window and the emission detection window, thetransmission self-contamination data representing radiation from thesecond radiation source undesirably detected by the first scintillationdetector.
 11. A method according to claim 10, further comprising usingthe first and second radiation sources and the first and secondscintillation detectors to perform a transmission scan of the object toacquire a set of transmission data; andusing the set of transmissionself-contamination data to correct the transmission data.
 12. A medicalimaging system, comprising:a first radiation detector and a secondradiation detector; a first radiation source for transmitting radiationto the first detector and not to the second detector; and a processingsystem coupled to control the first and second detectors and theradiation source, wherein the processing system is configured to controlthe first radiation source and the second detector to measure radiationtransmitted by the first source which is detected by the seconddetector.
 13. A medical imaging system according to claim 12, whereinthe processing system is further configured to:generate a set ofcalibration data corresponding to the radiation transmitted by the firstsource which is detected by the second detector: use the first andsecond detectors and the radiation source to perform a transmission scanof an object to be imaged to acquire a set of transmission data; andcorrect the transmission data with the calibration data.
 14. A nuclearmedicine imaging system, comprising:a first radiation detector and asecond radiation detector; a first radiation source and a secondradiation source for transmitting radiation to the first detector andnot to the second detector; and a processing system coupled to controlthe first and second detectors and the first and second radiationsources, wherein the processing system is configured to:control thefirst and second radiation sources and the first and second detectors toperform a nuclear medicine transmission scan of an object to acquire aset of transmission data; and use a second set of data to correct theset of transmission data, wherein the second set of data is based on acalibration scan, the calibration scan characterized by a measurement ofradiation transmitted by the first source which is detected by thesecond detector.
 15. A nuclear medicine imaging system according toclaim 14, wherein the set of transmission data comprises a plurality ofcounts, and wherein the processing system is further configured tosubtract counts from the set of transmission data based on the secondset of data.
 16. An apparatus for acquiring image data associated withan object, the apparatus comprising:a first radiation detector and asecond radiation detector; a first radiation source and a secondradiation source, wherein the first radiation source is for transmittingradiation to the first detector and not to the second detector; meansfor illuminating the second detector with radiation from the secondradiation source while maintaining the first radiation source in anon-transmitting state; means for detecting radiation from the secondradiation source with the first detector while the second detector isilluminated with radiation from the second radiation source; means forgenerating calibration data based on the radiation detected within thefirst detector, the calibration data for correcting a set of datarepresentative of the object.
 17. An apparatus according to claim 16,further comprising:means for performing a transmission scan of theobject to acquire a set of transmission data representative of anobject; and means for using the set of calibration data to correct theset of transmission data.
 18. An apparatus according to claim 17,wherein the set of transmission data comprises a plurality of counts,the gamma camera system further comprising means for subtracting countsfrom the set of transmission data based on the calibration data.